When blood or biological fluids contact polymeric biomaterials, several processes occur including adsorption of plasma proteins, platelet adhesion and activation, and activation of the immune complement and coagulation pathways. The ultimate consequences of these processes are fouling and failure of the device and serious clinical complications. Modification of the biomaterial surface with albumin, heparin, and water soluble synthetic polymers such as poly(ethylene oxide) (PEO) is known to decrease or prevent these complications [1-13]. Surface modification of biomaterials with those surface-passivating molecules, however, has been rather difficult, since most of the biomaterials do not have any chemically active functional groups on their surfaces.
With respect to modifying surfaces with surface-passivating polymer molecules such as albumin and PEO, the simplest method is physical adsorption [5,8,14-16]. To utilize this method a hydrophobic surface, such as polypropylene or dimethyldichlorosilane (DDS)-glass, is exposed to an aqueous solution of albumin or PEO. Polymer molecules spontaneously adsorb to the surface due to the hydrophobic interaction. There is no chemical reaction between the polymer and surface, just physical interaction. The polymer layer is therefore not permanently bound to the surface, and generally can be removed with appropriate washing. This method is often ineffective for long-term prevention of protein adsorption and platelet adhesion, as other molecules or cells with a greater affinity for the surface may displace the physically adsorbed polymer. Furthermore, it has been suggested that drying of the surface may result in a heterogeneous polymer layer due to polymer molecules migrating with the receding water during the drying process. Drying and rehydration may also result in removal of the adsorbed polymer molecules from the surface. The surface concentration of the adsorbed albumin on DDS-glass was reduced by more than 15% by simple drying and rehydration [17]. Simple adsorption is therefore not appropriate for permanent applications of the modified surface.
For the covalent grafting of the surface-passivating molecules, the biomaterial surfaces have been premodified by various methods such as simple polymer adsorption, chemical surface-polymer coupling, and graft polymerization [1-8]. Previous approaches to permanently graft hydrophilic polymers to surfaces fall into two categories: graft polymerization and graft coupling. In graft polymerization, polymer chains are synthesized from the reactive surface. Alternatively, graft coupling binds polymer molecules to the surface through chemical reactions. Plasma polymerization and graft polymerization are the methods commonly employed in the former category, while various chemical methods comprise the latter approach.
Plasma polymerization, also known as glow discharge polymerization, is a method of polymerizing monomers from the vapor phase at low pressure. It is commonly used to polymerize monomers onto surfaces, resulting in a highly crosslinked polymer layer [18]. These highly cross-linked polymer layers are generally unable to prevent platelet adhesion on surfaces [6,19,20]. Rather, single chains of grafted polymer are necessary to exert the surface passivating effects. Graft polymerization is another process by which polymers are synthesized directly onto a reactive surface. In this approach, the substrate to be grafted is placed in a solution of the monomer. Polymerization is initiated by exogenous energy sources such as UV light, heat, or .gamma.-radiation, or by chemical initiators [21-25]. The polymerization reaction proceeds both in the bulk solution yielding free-floating polymer chains and from the reactive sites on the surface yielding grafted chains. The surface reactive groups may resemble the monomer, so that the surface is included into a growing polymer chain, or it may otherwise contain the initiator groups, so that the grafted polymer chain actually begins growing from the surface. Trichlorovinylsilane (TCVS) has been used to modify surfaces prior to graft polymerization [26]. Producing consistent grafted surfaces is difficult to achieve with the graft polymerization technique. Since the polymers are synthesized in situ, there is little if any control over the degree of polymerization and the polymer grafting efficiency (i.e., the number of grafted polymer chains per unit area of surface). Polydispersity in molecular weights of the grafted chains confounds surface characterization. Depending on the polymerization method crosslinking of polymer chains may occur, leading to a 3-dimensional network of grafted polymer on the surface, rather than a layer of grafted individual chains. As mentioned above, the surface passivating effects of surface bound polymers is attributable to free polymer chains, not crosslinked polymer gels.
Chemical methods to graft purified polymer chains to surfaces have also been widely used [1-4, 27-32]. These methods employ chemical reactions between polymer molecules and reactive sites on the surface. Chemical grafting therefore relies on the presence of complementary reactive groups on the polymer and the surface, necessitating a different specific approach to each system. To utilize a specific polymer, it must already have a suitable reactive group or must be modified to contain such a reactive group. Because of this limitation chemical methods are of limited utility for generalized application. Each polymer-surface system is essentially a completely new project to be optimized for polymer synthesis and grafting conditions. The chemical methods have been used for immobilizing enzymes on supports and grafting polymer chains to various surfaces. Recently, surfaces have been modified with prefunctionalized polymers which are activatable by UV light, heat, or .gamma. radiation. The use of activatable polymers allow grafting of polymer chains to otherwise inert surfaces. Examples of such polymers and their utility in surface modifications to prevent protein adsorption and cell adhesion are in the literature [9-11,13,33]. .gamma.-radiation has also been used to graft hydrophilic polymer chains to various solid polymeric surfaces. In our laboratory we have synthesized several polymers which are activated by .gamma.-radiation [17,34-36]. When activated, the polymer chains react with the surface and form a chemically bound polymer layer. In this method the polymer molecules adsorb from solution to the substrate surface and are then activated with .gamma.-radiation to permanently graft to the surface. This approach requires that the highly reactive intermediates be in close proximity to the surface so that the activated polymer chains react with the surface. If the reactive groups are oriented away from the surface, they may react with adjacent polymer chains or with solvent water molecules. This results in very low grafting efficiency.
While hydrophilic polymer chains can be graft coupled to solid polymeric surfaces rather easily, their grafting to non-organic surfaces such as metals and glasses has been difficult. In contrast, the use of hydrophobic alkyl side chains to form ultrathin self-assembled polymeric films on solid surfaces has been done. [44-48]. Metals such as titanium and aluminum used in biomedical devices face the same blood compatibility problems as polymeric materials. The inorganic nature of metal surfaces makes them particularly difficult to graft with hydrophilic polymers such as PEO. Plasma polymerization has been used to graft metal surfaces with synthetic hydrophilic polymers [18]. As previously stated, however, the high degree of crosslinking inherent in plasma polymerization processes is not suitable for preparing protein and platelet resistant surfaces.
Recently, simple methods for the covalent grafting of surface-passivating molecules such as albumin and PEO have been developed [9-11]. The use of ultraviolet (UV) light, heat, or .gamma. radiation for the covalent grafting of albumin or PEO to chemically inert surfaces such as polyethylene, polypropylene, polycarbonate, poly(vinyl chloride), and dimethyldichlorosilane (DDS)-coated glass was explored. In addition to albumin and PEO, poly(ethylene oxide)/poly(propylene oxide)/poly(ethylene oxide) (PEO/PPO/PEO) triblock copolymers were grafted using .gamma.-irradiation [13].
Prior art methods have all rendered the modified surfaces resistant to protein adsorption and platelet adhesion. However, while these approaches are simple, highly effective, and can be applied to fully assembled devices in various shapes, they require introduction of activatable groups to the surface-passivating molecules. The preparation of activatable albumin or PEO requires rigorous synthesis and purification chemistry. In addition, these methods can only be used on polymeric biomaterials. Grafting of surface-passivating molecules to metal, glass, or ceramic surfaces is difficult with these methods.
Metallic materials are extensively used for construction of long term implantable cardiovascular devices such as prosthetic heart valves and stents. Several metallic materials enjoy wide acceptance due to their in vivo corrosion resistance. The electrochemical activity of blood makes corrosion resistance a major concern. As these materials are in contact with blood, it is desirable to improve their biocompatibility.
Titanium is used exclusively in prosthetic heart valves. Cobalt-chrome alloys, (such as STELLITE.TM. 21 and HAYNEST.TM. 25) tantalum, and nickel alloys are also used in prosthetic heart valves. These materials exhibit excellent corrosion resistance due to the high stability of the oxide layers on their surfaces. This oxide layer protects the deeper material from further oxidation. Clean titanium forms a tenacious titanium oxide layer, which is stable to saline solution, when exposed to air due to the high reactivity of the metal. Aluminum also reacts similarly, but the oxide layer which is formed is not stable to saline solution. The process of anodizing is required to stabilize aluminum oxide layers.
Blood contacting metallic materials are also found in cardiopulmonary surgical devices such as blood oxygenators. The heat exchangers which control blood temperature during surgery have very large surface areas for efficient heat transfer. They are made of anodized aluminum or stainless steel. Neither material is considered blood compatible, so the anticoagulant heparin is used to prevent blood clotting. Rendering the surfaces blood compatible by surface passivation may greatly reduce the doses of heparin required during such surgeries.